Many bioactive molecules with therapeutic potential for the central nervous system (CNS) exert wide-ranging activities that can lead to unwanted side effects if they are delivered systemically. A number of potential therapies for neurological conditions have failed in clinical trials for this reason. Site-specific delivery of potential therapies is increasingly recognized as an important goal for many neurological conditions, and biomaterial vehicles represent a promising means of achieving this goal. However, suitable vehicles are not presently available for the restricted targeting of growth factors or other diffusible molecules, or the placement of extracellular matrices for axon growth or cell migration. Furthermore, neural stem/progenitor cell grafts may hold promise for replacing lost cells in certain neurological conditions, but most current grafting procedures are regarded as resulting in suboptimal survival and differentiation of cells. Biomaterial vehicles may be able to improve grafting efficiency by providing grafted cells with support matrix and molecular substrates that help them overcome the shock of the grafting procedure and integrate and differentiate better in to host tissue.
Injectable biomaterials represent a rapidly advancing new area for delivery of therapeutic molecules. The U.S. F.D.A. has already approved certain biomaterials for slow release, long-term systemic drug delivery, while others are currently undergoing clinical trials The basic precedent of using biomaterials for drug delivery in humans has thus been set, but no materials are currently approved for delivery into the CNS.
Hydrogels are a class of materials that have significant promise for use in soft tissue and bone engineering (Lee et al. (2001) Hydrogels for tissue engineering, Chemical Reviews 101, 1869-1879). An important feature of hydrogels that makes them attractive for these applications is their well hydrated, porous structure that can mimic natural extracellular matricies (Peppas et al. (2000) Physicochemical foundations and structural design of hydrogels in medicine and biology, Annual Review of Biomedical Engineering 2, 9-29). To replace natural materials, however, many structural and functional features must be built into synthetic hydrogels. Desirable features include: biocompatability; degradability to allow cell ingrowth; injectable yet also fast setting in the wound site; mechanical properties that can be tuned for different uses; control over cell adhesion to the hydrogel matrix; and tunable sustained release of growth factors and other biologically active agents. There are many examples where some, or even most of these features have been incorporated into hydrogels. However, in many cases, hydrogel synthesis and formation becomes very complicated, which limits the practicality of such materials. More importantly, the complexity of these systems, combined with limited means for adjustment of molecular parameters, leads to the inability to independently adjust most of the features. For example, it would be advantageous to be able to adjust scaffold rigidity while maintaining a constant hydrogel mesh size. Such a system would allow one to directly measure the effects of scaffold rigidity on cell proliferation. Also, since hydrogel degradation is commonly accomplished using degradable crosslinkers (e.g. in PEG based hydrogels), it can be difficult to adjust degradation rate without also altering crosslink density, and hence initial gel mechanical properties. It would be desirable to have a hydrogel system where many of these desirable adjustable features (e.g. gel strength, gel density, adhesive capability, degradation rate, growth factor release rate) could be controlled independently so that, e.g., meaningful evaluations of their roles in tissue regeneration could be systematically evaluated. Currently, in many systems it is difficult to identify the most important gel characteristics, since many features are adjusted simultaneously.
Current hydrogel technology utilizes both naturally-derived macromolecules and synthetic polymers. Generally, hydrogels prepared from natural polymers possess desired biological signalling capability but may lack desired material properties, e.g. low sample rigidity, and may also be problematic due to immunogenicity and pathogen transmission issues. By contrast, synthetic polymers can be engineered for desired material properties but may display limited cytocompatibility. One approach to increase the cytocompatibility of synthetic polymers is to incorporate peptide epitopes, for example RGD motifs. However, incorporating these motifs into preformed polymers in a regiospecifically controlled manner is extremely difficult. In addition, these scaffolds are structurally homogeneous (not porous) on the microscale due to their underlying molecular network structure, which can limit cell proliferation. These systems must undergo additional processing (e.g. freeze-thaw cycling, particulate leaching, microsphere sintering and non-woven fiber formation) in order to introduce microscale porosity in the gel network. However, despite their dilute, porous nature, these well hydrated materials must also be mechanically rigid. This apparent contradiction, rigidity from a dilute porous scaffold, must be inherently addressed by the design of constituent molecular crosslinks (chemical and/or physical) formed during the hydrogelation process. However, introducing chemical crosslinks may be biologically problematic since by-products from the crosslinking chemistry may be toxic and difficult to remove from the scaffold. It would be desirable to generate benign, biocompatible chemical or physical crosslinking methods for either in vitro gelation for eventual incorporation in the body or direct, rapid in vivo gelation. An additional design complication is that hydrogel rigidity seemingly precludes any viable processability in preformed scaffolds. For example, one may wish to form a rigid tissue engineering construct in vitro but subsequently inject it into a host for tissue regeneration. Injection is not possible in a permanently crosslinked, rigid network. In short, the many seemingly contradictory features required in hydrogels for tissue engineering applications severely prohibits the use of materials with a large degree of adjustability in their properties. The polymers used in most synthetic hydrogels simply do not contain enough functionality to allow tuning of degradability, adhesion or gel strength without compromising other necessary properties.
Amphiphilic diblock copolypeptide hydrogels (referred to herein as “DCH”) are synthetic materials with many features that make them attractive as tissue engineering candidates for applications that are likely to require progressive adjustment and fine-tuning of material properties (Pakstis et al. (2004) Effects of Chemistry and Morphology on the Bifunctionality of Self-Assembling Diblock Coplypeptide Hydrogels, Biomacromolecules 5, 312-318); Deming (2005) Polypeptide hydrogels via a unique assembly mechanism, Soft Matter 1, 28-39). The present inventors have previously used a combination of chemical synthesis and structural characterization to establish an understanding of DCH structure-property relationships that allows a high level of control over gel strength, gel porosity, gel functionality and media stability; and many of these properties can be adjusted independently of each other (Nowak et al. (2002) Rapidly recovering hydrogel scaffolds from self-assembling diblock copolypeptide amphiphiles, Nature 417, 424-428; Breedveld et al. (2004) Rheology of block copolypeptide solutions: hydrogels with tunable properties, Macromolecules, 3943-395)). DCH are physically associated gels that can be deformed and thinned by stress and injected through small-bore cannulae, after which they rapidly re-assemble into rigid gel networks (Nowak et al. (2002) (supra)). These properties provide DCH with the potential for facile and minimally invasive delivery in vivo. DCH form elastic gels with fibril-like nanostructures and porous microstructures theoretically suitable for integration with host cells (Nowak et al. (2002), (supra); Deming (2005), (supra).
It was unknown prior to the present invention whether DCH could be generated that are suitable for administration to the CNS, or whether such DCH could serve as depots for biologically active materials or act as scaffolds to support cell migration in the CNS.